1. Field of the Invention
This invention relates generally to optical coherence tomography (OCT) systems for three-dimensional (3D) imaging and more particularly to a power-efficient OCT system and method for rapid 3D ocular imaging without motion artifacts.
2. Description of the Related Art
The optical coherence tomography (OCT) art has evolved over time from the optical time-domain reflectometry (OTDR) art described by Barnoski et al. in 1976 [J. K. Barnoski, S. M. Jensen, “Fiber waveguides: A novel technique for investigation attenuation characteristics,” Appl. Opt., 15, 2112-5 (1976)]. OTDR was first employed to measure the elapsed time (t) and intensity of light reflected along a SINGLE path in optical fiber to determine the distance (d=ct) to problems along the fiber such as attenuation and breaks, making it a useful tool in optical network trouble-shooting. The original idea of OTDR consists in launching a rather short and high power optical impulse into the tested fiber and a consequent incoherent detection of optical power backscattered along the z-axis of the fiber as a response to the test impulse. The detected signal provides the detailed picture of the local loss distribution along the fiber caused by any of the attenuation mechanisms or some other nonhomogeneities on the fiber. In the same year, Kompfner et al. [. Kompfner and H. Park, Int. J. Electron., 41, 317 (1976)] proposed a system for the coherent detection of a series of such backscattered pulses to see through otherwise opaque material.
Several years later, in 1981, Park et al. [H. Park, M. Chodorow, and R. Kompfner, “High Resolution Optical Ranging System,” Appl. Opt., 20, 2389-94 (July 1981)] reported experimental results for the Kompfner et al proposal, which adapts the incoherent OTDR technique by splitting the short and powerful optical impulse signal into TWO physical channels and combining the optical power backscattered from a reference mirror with that backscattered from a test sample. Using coherent detection, Park et al. were able to measure test sample reflections from the particular test sample z-axis locus corresponding to the reference mirror position on the z-axis. Axial motion of the reference mirror serves to move the test sample reflection detection locus along the z-axis. This OTDR technique was denominated coherent OTDR (CO-OTDR) by some practitioners. Park et al. were able to achieve 1.7 mm z-axis resolution and proposed the addition of two-dimensional (2D) scanning means to permit three-dimensional imaging in an otherwise opaque test sample. This proposal may be properly denominated “tomography,” which denotes “an imaging technique using sections or planes to visualize the interior” of a test sample.
Thereafter, in 1987, Youngquist et al. [R. C. Youngquist, S. Carr, and D. E. N. Davies, “Optical coherence-domain reflectometry: a new optical evaluation technique,” Opt. Lett., 12, 158–160 (March 1987)] first proposed a modification of the CO-OTDR technique using a continuous wave optical source signal having a short coherence length. They denominated their incoherent source method “optical coherence-domain reflectometry (OCDR).” The OCT denomination later appeared in the art [e.g., Huang et al., “Optical Coherence Tomography,” Science, 254, 1178–81 (November 1991)]. With OCDR, the output signal from an incoherent optical source is said to have a “short coherence length” when its autocorrelation function has a single peak that is relatively narrow in time. The OCDR method splits the incoherent signal into two channels and combines the reflections from a reference mirror and a test sample at a detector where the signals interfere to form fringes whose intensity represent the reflectance from a volumetric region of the test sample at a position on the z-axis defined by the reference mirror position, the extent of which is defined in the xy plane by a focus area and in depth on the z-axis by the signal coherence length. A transverse scanner can be added to map this reflectance over a transverse “slice” of the test sample in the xy plane. Mapping a series of these xy slices by moving the reference mirror along the z-axis results in a three-dimensional image (tomography) of an internal test sample volume (3D-OCT).
Since 1987, numerous practitioners have proposed improvements to the OCT art, many of which are discussed in a review by Fercher et al. [A. F. Fercher, C. K. Hitzenberger, “Optical coherence tomography,” Chapter 4 in Progress in Optics 44, Elsevier Science B.V. (2002)]. Although the OCT art offers many advantages for biological tissue mapping, especially in the eye, many practical problems have been identified over the years, such as those relating to the design of practical interferometric scanning and detection systems, generation of partially coherent fields, improved detection (scan) speeds, and the elimination of movement artifacts arising from involuntary eye movement during scanning.
FIG. 1 illustrates a typical OCT scanner 12 from the prior art. The interferometer 13 splits a signal S from a broadband source 14 into a reference signal SR and a sample signal SS. The reference signal SR is directed to a reference reflector 16 disposed to move in either direction along the z-axis and the sample signal SS focuses through the scanning optics 18 and the objective lens 20 to some point 22 within the test sample 24 under test (e.g., tissue). After scattering back from point 22 in test sample 24, the modified sample field ES mixes with the reflected reference signal field ER on the surface of a photodetector 26. Assuming that photodetector 26 captures half of the light from the reference and sample arms of interferometer 13 (the other half returns to the source if a normal 50:50 beam splitter is used), the signal intensity impinging on photodetector 26 isID=(|SD|2)=0.5(IR+IS)+Re{ER*(t+τ) ES(t)}≡IDC+IPIX  [Eqn 1]where IR and IS are the mean (DC) intensities of the reflected signals returning from the reference and sample arms of the interferometer. The second term in Eqn. 1 is the cross-correlation signal IPIX, which depends on the optical time delay τ established by the z-axis position of reference reflector 16 and represents the amplitude of the interference fringes that carry information about the structure of point 22 in test sample 24; the envelope of this fringe signal may correspond to a single “pixel” in a 2D image of test sample 24. The presence and nature of these interference fringes depends on the alignment of the temporal and spatial characteristics of the reflected fields ES and ER. Thus, interferometer 13 functions as a cross-correlator and the amplitude IPIX of interference signal generated after integration on the surface of detector 26 provides a measure of the cross-correlation amplitude. The z-axis thickness of point 22 depends on the coherence length of broadband signal S. Various techniques are known in the art for modulating τ (e.g. by vibrating reflector 16) to facilitate separation of the cross-correlation signal IPIX from the mean component IDC of intensity ID at detector 26 (the first term of Eqn. 1). When IDC greatly exceeds IPIX, the detector may be operating in the excess intensity noise regime where the effective signal-to-noise ratio (SNR) is degraded. Movement of mirror 16 along the z-axis facilitates measurement of reflectance from test sample 24 at numerous points along the z-axis. Scanning optics 18 may be arranged to facilitate scanning of en face images over a 2D (xy) plane within test sample 24 at various (usually sequential) z-axis locations.
As shown in FIG. 2 from the prior art, any of several scanning patterns may be used to obtain three-dimensional (3D) image data sets with OCT scanner 12. Most practitioners refer to a longitudinal imaging procedure wherein the longitudinal scan lines directed along the z-axis in the image correspond to A-scans and the transverse scan along the x-axis in FIG. 2 (or the y-axis) advances at a slower pace to build the B-scan image 28 illustrated at the top of FIG. 2. This may be reversed so that the transverse scanner produces the fast lines in the image and the longitudinal scanner advances more slowly to build the B-scan image 30 illustrated in the middle of FIG. 2, which simplifies the production of transverse en face images for a fixed reference path, such as the en face image 32 illustrated in FIG. 1 bottom. A first transverse scanner scans the test sample along the lines (x-axis) in image 32 while a second transverse scanner advances more slowly along the second co-ordinate (y-axis) in image 32. A transverse slice (en face image) is thereby collected at each of several different depths on the z-axis, either by advancing the optical path difference in steps after each complete transverse scan or continuously at a speed for which the depth position of the point in the top left corner of the image and the depth position in the bottom right corner of the image do not differ by more than half the depth resolution. This provides one of the fastest methods for recording a 3D image data set for a region within sample 24.
Recording a single typical 3D image of, e.g., the retina of a human eye, requires at least one second of scanning time in the present art. Involuntary eye movements occurring during this recording period may introduce distortions into the 3D image data, and consequently, may distort and degrade any 2D diagnostic images derived from such data. For example, Podoleanu et al [A. Gh. Podoleanu, J. A. Rogers, D. A. Jackson, S. Dunne, “Three dimensional OCT images from retina and skin,” Opt. Express, 7, 292–298 (2000)] suggest that en face OCT images are preferred for reasons of speed but also prone to “blurring” arising from test sample motion. In the commonly-assigned U.S. Pat. No. 6,137,585, entirely incorporated herein by reference, Hitzenberger describes a differential OCT system in which artifacts arising from z-axial components of sample motion can be eliminated by using a reference reflector defined by the sample (e.g., the cornea in ocular OCT imaging) that moves along the z-axis with the sample. However, this method is not useful for eliminating artifacts arising from transverse (xy plane) components of sample motion generally.
Transverse motion artifacts are embodied as misalignment of sequential transverse slices (en face images) recorded at different sample depths and thus may be eliminated by detecting and aligning image features of sequential transverse images, provided these features are present in several sequential images. Because of the very narrow depth of each OCT image slice and the curvature of the retina, transverse OCT images of the retina have a fragmented appearance that makes it difficult to find common features in sequential images. This problem is well-appreciated in the art and has been addressed by several practitioners.
For retinal imaging, some have suggested that the motion artifact alignment problem can be resolved by recording, in parallel to each OCT slice, a separate image with the wider depth range needed to reveal test sample features sufficient to guide any realignment of the OCT slices necessary to remove motion artifacts. Such a second image may be obtained, e.g., by employing a separate detector operating in a scanning laser ophthalmoscope (SLO). In principle, the SLO images may reveal the precise timing and degree of any transverse eye motion after scan completion with the help of visible landmark features common in each of these images.
In U.S. Pat. No. 5,975,697, Podoleanu et al. describe an optical mapping apparatus for measuring en face images with adjustable depth to permit correction of the images for curvature of the retina at the back of the lens of the eye. Podoleanu et al. describe the many considerable difficulties with using OCT and SLO en face images in parallel and suggest elaborate procedures intended to eliminate some of these problems, including readjusting the SLO image depth resolution, recording OCT slices at several different resolutions, and employing common receiver optics for both OCT and SLO image channels. Podoleanu et al. suggest that their elaborate procedures, while slow, may permit the useful comparison of OCT retinal image data to existing SLO image databases for medical diagnosis. Disadvantageously, with this method, part of the source light power must be diverted to a separate SLO detector, decreasing the SNR of the OCT image channel. Podoleanu and Jackson [A. Gh. Podoleanu, D. A. Jackson, “Noise Analysis of a Combined Optical Coherence Tomograph and a Confocal Scanning Ophthalmoscope,” Appl. Optics, 38, 2116–7, April 1999] suggest that their OCT channel SNR must be traded off to permit the simultaneous acquisition of OCT and SLO en face images. They also note the speed penalty associated with this SNR degradation and with their method of combining OCT and SLO images of the retina. Moreover, this method disadvantageously requires an additional detector, amplifier, and frame grabber to avoid the detector SNR limitations encountered in the excess intensity noise dominated regime. Later, Rogers et al [J. A. Rogers, A. Gh. Podoleanu, G. M. Dobre, D. A. Jackson and F. W. Fiske, “Topography and volume measurements of the optic nerve using en-face optical coherence tomography,” Optics Express, 9, 533–45, 05 Nov. 2001] describe an application of the en-face OCT scanning technique to optic nerve topography. While Rogers et al. stated that the confocal channel was not absolutely necessary, it greatly helped to track the relative eye movements in the OCT en face images. For this purpose, Rogers et al. also require a separate detector and beam splitter to record their OCT signal and they observe that further study is needed to determine the optimum number of frames to be superposed to realize the best advantages of their suggested method. Their additional beam splitter diverts part of the available light away from the OCT receiver, which reduces the light power reaching the OCT detector via the sample arm and thereby reduces the sensitivity of the OCT detector channel.
More recently, Hitzenberger et al. [C. K. Hitzenberger, P. Trost, P. W. Lo, and Q. Zhou, “Three-dimensional imaging of the human retina by high-speed optical coherence tomography,” Opt. Express, 11, 2753–61 (October 2003)] suggest generation of SLO-like images by projection of the transversal OCT image slices on top of each other, thereby avoiding the necessity of the second or parallel SLO imaging channel suggested in earlier publications. The proposed SLO-like images do not require a second detector so the OCT channel sensitivity is unaffected thereby but these SLO-like images are still somewhat distorted by movement artifacts, and therefore cannot be used to re-align 3D OCT image data.
Useful solutions to the OCT motion artifact problem are limited by several well-known OCT system noise problems. OCT systems like OCT scanner 12 illustrated in FIG. 1 (discussed above) are subject to three major noise sources; receiver-amplifier noise; shot noise: and excess intensity noise. Receiver noise dominates in the regime where the light power ID (Eqn. 1) available at the detector is very low. The receiver noise dominated regime can usually be avoided by using state-of-the-art electronics and sufficient optical source power. When IDC greatly exceeds IPIX, (Eqn. 1) the detector may enter the excess intensity noise regime where the effective signal-to-noise ratio (SNR) is degraded. Excess intensity noise dominates in the regime where the light power ID at the detector is very high so that more light power does not improve effective sample SNR at the detector and may instead reduce SNR if the additional light power consists only of the IDC term. To avoid the excess intensity noise dominated regime, the reference light intensity IR is usually attenuated, typically by a factor of 100 or more, to reduce IDC with respect to IPIX. Shot noise arises from the inherent quantum nature of light and cannot be avoided, so it dominates in the intermediate regime between the receiver noise and excess intensity noise regimes. However, in the shot noise regime, sensitivity improves linearly with the light power IS′ backscattered by the sample. For a given source power, the optimum OCT system sensitivity is achieved when operating in the shot noise dominated regime but this condition limits the usefulness of the available source power, most of which must be discarded to avoid the excess intensity noise regime.
Useful OCT scanning speed depends on the available OCT detector channel sensitivity. The OCT detector sensitivity problem includes the excess intensity noise issue mentioned by Podoleanu and Jackson (above) and also other issues, such as the polarization distortion problem discussed, for example, in U.S. Pat. No. 6,134,003 issued to Tearney et al., who suggest using Faraday rotators or optical circulators in a fiber optic OCT apparatus to improve OCT system sensitivity. Similar OCT system designs with improved sensitivity, based on optical circulators, have been suggested for high speed imaging applications by Rollins et al. [A. M. Rollins and J. A. Izatt, “Optimal interferometer designs for optical coherence tomography,” Opt. Lett. 24 1484–6, November 1999] and in U.S. Pat. No. 6,657,727 issued to Izatt et al. Elsewhere, in U.S. Pat. No. 6,615,072, Izatt et al. also suggest using a polarizing element, such as a Faraday rotator or optical circulator, on the optical path to compensate for variations in interference intensity at the detector caused by variation in fiber birefringence in a power effective fiber optic OCT probe apparatus. Similarly, in U.S. Pat. No. 6,385,358, Everett et al. describe a birefringence insensitive OCDR system that uses a Faraday rotator to cancel polarization mismatch arising from the use of inexpensive disposable non-polarization maintaining optical fiber in the sample arm, thereby permitting its use in various disposable clinical devices such as catheters, guidewires, and hand-held instruments or probes. In U.S. Pat. No. 5,202,745, Sorin et al. discuss an OCDR system that employs polarization diversity signal processing methods to overcome the effects on OCT detection sensitivity of the polarization distortion usually found in optical fibers and other system components. Disadvantageously, for some applications, polarizing elements such as Faraday rotators and optical circulators may be too expensive with respect to simple polarizing elements and retardation plates. Further, the optical circulator is presently only available for the wavelength range of 1300–1550 nm, and not for the 800 nm region preferred for retinal OCT.
The typical Michelson interferometer splits the source beam power equally into a sample arm signal SS and a reference arm signal SR (FIG. 1). After reflection of the light from the sample, 50% of the reflected sample light is directed to the detector and 50% toward the source. Therefore, with a sample reflectivity R, only 0.5×0.5×R=0.25×R of the source light power reaches the detector by way of the sample. A similar fraction of 0.25 of the emitted light power reaches the detector by way of the reference arm, assuming a reference mirror reflectivity of 100%. Disadvantageously, the reference power IR (Eqn. 1) must often then be further attenuated to avoid the excess intensity noise dominated regime at the detector. For example, Hoeling et al. [B. Hoeling, A. Fernandez, R. Haskell, E. Huang, W. Myers, D. Petersen, S. Ungersma, R. Wang, M. Williams and S. Fraser, “An optical coherence microscope for 3-dimensional imaging in developmental biology,” Opt. Express, 6, 136–145 (2000)] suggest reducing the reference power by 75% to improve detector SNR by 40% by avoiding the excess intensity noise regime.
There is accordingly a clearly felt need in the art for an OCT imaging technique that can inexpensively resolve these test sample motion artifact and detector sensitivity problems in a manner that reduces the acquisition time for accurate 3D OCT images of biological tissues, such as the retina. These unresolved problems and deficiencies are clearly felt in the art and are solved by this invention in the manner described below.